Implantable stimulation devices generate and deliver electrical stimuli to nerves and tissues for the therapy of various biological disorders, such as pacemakers to treat cardiac arrhythmia, defibrillators to treat cardiac fibrillation, cochlear stimulators to treat deafness, retinal stimulators to treat blindness, muscle stimulators to produce coordinated limb movement, spinal cord stimulators to treat chronic pain, cortical and deep brain stimulators to treat motor and psychological disorders, occipital nerve stimulators to treat migraine headaches, and other neural stimulators to treat urinary incontinence, sleep apnea, shoulder sublaxation, etc. The present invention may find applicability in all such applications and in other implantable medical device systems, although the description that follows will generally focus on the use of the invention in a Bion™ microstimulator device system of the type disclosed in U.S. Patent Application Publication 2010/0268309.
Microstimulator devices typically comprise a small generally-cylindrical housing which carries electrodes for producing a desired stimulation current. Devices of this type are implanted proximate to the target tissue to allow the stimulation current to stimulate the target tissue to provide therapy for a wide variety of conditions and disorders. A microstimulator usually includes or carries stimulating electrodes intended to contact the patient's tissue, but may also have electrodes coupled to the body of the device via a lead or leads. A microstimulator may have two or more electrodes. Microstimulators benefit from simplicity. Because of their small size, the microstimulator can be directly implanted at a site requiring patient therapy.
FIG. 1 illustrates an exemplary implantable microstimulator 100. As shown, the microstimulator 100 includes a power source 145 such as a battery, a programmable memory 146, electrical circuitry 144, and a coil 147. These components are housed within a capsule 202, which is usually a thin, elongated cylinder, but may also be any other shape as determined by the structure of the desired target tissue, the method of implantation, the size and location of the power source 145 and/or the number and arrangement of external electrodes 142. In some embodiments, the volume of the capsule 202 is substantially equal to or less than three cubic centimeters.
The battery 145 supplies power to the various components within the microstimulator 100, such the electrical circuitry 144 and the coil 147. The battery 145 also provides power for therapeutic stimulation current sourced or sunk from the electrodes 142. The power source 145 may be a primary battery, a rechargeable battery, a capacitor, or any other suitable power source. Systems and methods for charging a rechargeable battery 145 will be described further below.
The coil 147 is configured to receive and/or emit a magnetic field that is used to communicate with, or receive power from, one or more external devices that support the implanted microstimulator 100, examples of which will be described below. Such communication and/or power transfer may be transcutaneous as is well known.
The programmable memory 146 is used at least in part for storing one or more sets of data, including electrical stimulation parameters that are safe and efficacious for a particular medical condition and/or for a particular patient. Electrical stimulation parameters control various parameters of the stimulation current applied to a target tissue including, but not limited to, the frequency, pulse width, amplitude, burst pattern (e.g., burst on time and burst off time), duty cycle or burst repeat interval, ramp on time and ramp off time of the stimulation current, etc.
The illustrated microstimulator 100 includes electrodes 142-1 and 142-2 on the exterior of the capsule 202. The electrodes 142 may be disposed at either end of the capsule 202 as illustrated, or placed along the length of the capsule. There may also be more than two electrodes arranged in an array along the length of the capsule. One of the electrodes 142 may be designated as a stimulating electrode, with the other acting as an indifferent electrode (reference node) used to complete a stimulation circuit, producing monopolar stimulation. Or, one electrode may act as a cathode while the other acts as an anode, producing bipolar stimulation. Electrodes 142 may alternatively be located at the ends of short, flexible leads. The use of such leads permits, among other things, electrical stimulation to be directed to targeted tissue(s) a short distance from the surgical fixation of the bulk of the device 100.
The electrical circuitry 144 produces the electrical stimulation pulses that are delivered to the target nerve via the electrodes 142. The electrical circuitry 144 may include one or more microprocessors or microcontrollers configured to decode stimulation parameters from memory 146 and generate the corresponding stimulation pulses. The electrical circuitry 144 will generally also include other circuitry such as the current source circuitry, the transmission and receiver circuitry coupled to coil 147, electrode output capacitors, etc.
The external surfaces of the microstimulator 100 are preferably composed of biocompatible materials. For example, the capsule 202 may be made of glass, ceramic, metal, or any other material that provides a hermetic package that excludes water but permits passage of the magnetic fields used to transmit data and/or power. The electrodes 142 may be made of a noble or refractory metal or compound, such as platinum, iridium, tantalum, titanium, titanium nitride, niobium or alloys of any of these, to avoid corrosion or electrolysis which could damage the surrounding tissues and the device.
The microstimulator 100 may also include one or more infusion outlets 201, which facilitate the infusion of one or more drugs into the target tissue. Alternatively, catheters may be coupled to the infusion outlets 201 to deliver the drug therapy to target tissue some distance from the body of the microstimulator 100. If the microstimulator 100 is configured to provide a drug stimulation using infusion outlets 201, the microstimulator 100 may also include a pump 149 that is configured to store and dispense the one or more drugs.
Turning to FIG. 2, the microstimulator 100 is illustrated as implanted in a patient 150, and further shown are various external components that may be used to support the implanted microstimulator 100. An external controller 155 may be used to program and test the microstimulator 100 via communication link 156. Such link 156 is generally a two-way link, such that the microstimulator 100 can report its status or various other parameters to the external controller 155. Communication on link 156 occurs via magnetic inductive coupling. Thus, when data is to be sent from the external controller 155 to the microstimulator 100, a coil 158 in the external controller 155 is excited to produce a magnetic field that comprises the link 156, which magnetic field is detected at the coil 147 in the microstimulator. Likewise, when data is to be sent from the microstimulator 100 to the external controller 155, the coil 147 is excited to produce a magnetic field that comprises the link 156, which magnetic field is detected at the coil 158 in the external controller. Typically, the magnetic field is modulated, for example with Frequency Shift Keying (FSK) modulation or the like, to encode the data.
An external charger 151 provides power used to recharge the battery 145 (FIG. 1). Such power transfer occurs by energizing the coil 157 in the external charger 151, which produces a magnetic field comprising link 152. This magnetic field 152 energizes the coil 147 through the patient 150′s tissue, and which is rectified, filtered, and used to recharge the battery 145 as explained further below. Link 152, like link 156, can be bidirectional to allow the microstimulator 100 to report status information back to the external charger 151. For example, once the circuitry 144 in the microstimulator 100 detects that the power source 145 is fully charged, the coil 147 can signal that fact back to the external charger 151 so that charging can cease. Charging can occur at convenient intervals for the patient 150, such as every night.
FIG. 3 illustrates salient portions of the microstimulator's power circuitry 160. Charging energy (i.e., the magnetic charging field) is received at coil 147 via link 152. The coil 147 in combination with capacitor 162 comprises a resonant circuit, or tank circuit, which produces an AC voltage at Va. This AC voltage is rectified by rectifier circuitry 164, which can comprise a well-known 4-diode bridge circuit, although it is shown in FIG. 3 as a single diode for simplicity. Capacitor 166 assists to filter the signal at node Vb, such that Vb is essentially a DC voltage, although perhaps having a negligible ripple. Intervening between Vb and the rechargeable battery 145 is charging circuitry 170, which ultimately takes the DC voltage Vb and uses it to produce a controlled battery charging current, Ibat. Charging circuitry 170 is well known. One skilled in the art will recognize that the power circuitry 160 may include other components not shown for simplicity.
It is generally desirable to charge the battery 145 as quickly as possible to minimize inconvenience to the patient. One way to decrease charging time is to increase the strength of the magnetic charging field by increasing the excitation current in the coil 157 of the external charger. Increasing the charging field will increase the current/voltage induced in the coil 147 of the microstimulator 100, which increases the battery charging current, Ibat. However, the strength of the magnetic charging field can only be increased so far before implant heating becomes a concern. One skilled in the art will understand that implant heating is an inevitable side effect of charging using magnetic fields. Heating can result from several different sources, such as eddy currents in conductive portions of the implant, or heating of the various components in the power circuitry 160. Implant heating is a serious safety concern; if an implant exceeds a given safe temperature (e.g., 41° C.), the tissue surrounding the implant may be aggravated or damaged.
The art has recognized that heating can be controlled by controlling the intensity of the magnetic charging field produced at the external charger 151. For example, the current flowing through charging coil 157 can be reduced to reduce the temperature of the implant during charging. The art has also recognized that heating can be regulated by duty cycling the charging field, i.e., by turning the charging field at the external charger 151 on and off. FIG. 4 generally shows the temperature of the implant, T(IPG), for two different duty cycles, DC1 and DC2, for a given magnetic charging field. The first duty cycle, DC1, equals 50%, because the magnetic charging field is on for 50% of the time (i.e., t1(on)=t1(off)). The second duty cycle, DC2, equals 75%, and hence the magnetic charger field stays on that much longer (i.e., t2(on)=3t2(off)). As one would expect, higher duty cycles result in higher temperatures in the implant: i.e., T1(IPG)<T2(IPG) as shown.
While changing the intensity or duty cycling of the magnetic charging field produced by the external charger 151 can be an effective means of controlling implant temperature, the inventors have realized that such approaches do not adequately address important issues. First, known prior approaches do not address whether the magnetic charging field intensity, duty cycle, or both, should be modified as a means of temperature control. Moreover, such prior techniques are not understood to consider efficient charging of the implant battery 145. Thus, one can change the intensity and/or duty cycle of the magnetic charging field to arrive at suitable temperature control, but the particular parameters chosen may provide a charging power to the battery that is unnecessarily low, which would prolong charging. Prolonged charging is inefficient, because that patient must wait an inordinate amount of time to fully charge the battery 145 in his or her implant. Understandably, patients do not desire charging to take any longer than necessary.
Finding optimal charging conditions (intensity, duty cycle) thus remains unknown with such prior art techniques, and this disclosure presents a technique to combat this problem, and to make charging more efficient from both a time and implant heating perspective.